Moving forward: A review of continuous kinetics and kinematics during wheelchair and handcycling propulsion

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Moving forward: A review of continuous kinetics and kinematics during wheelchair
and handcycling propulsion
Kellie M. Hallorana, Michael Fochta, Alexander Teagueb, Joseph Petersc, Ian Ricec, Mariana E. Kersha,b,d,∗
aDepartment of Mechanical Science and Engineering, University of Illinois Urbana-Champaign
bCarle Illinois College of Medicine, University of Il linois Urbana-Champaign
cDepartment of Kinesiology, University of Il linois Urbana-Champaign
dBeckman Institute for Advanced Science and Technology, University of Illinois Urbana-Champaign
Abstract
Wheelchair users (WCUs) face high rates of upper body overuse injuries, especially in the shoulder. As exercise
is recommended to reduce the high rates of cardiovascular disease among WCUs, it is becoming increasingly
important to understand the mechanisms behind shoulder soft-tissue injury in WCUs to help prevent future
injuries. Understanding the kinetics and kinematics during upper-limb propulsion in wheelchair users is the
first step toward evaluating soft-tissue injury risk during both everyday and athletic propulsion modes. This
paper examines continuous kinetic and kinematic data available in the literature for four common propulsion
modes. Two everyday modes (everyday wheelchair use and attach-unit handcycling) are examined, as well
as two athletic modes (wheelchair racing and recumbent handcycling). These athletic modes are important
to characterize, especially considering the higher contact forces, speed, and power outputs often experienced
during these athletic propulsion modes that could be putting users at increased risk of injury. Understanding
the underlying kinetics and kinematics during various propulsion modes can lend insight into shoulder loading,
and therefore injury risk, during these activities and inform future exercise guidelines and programs for WCUs.
Keywords: exercise, physical activity, racing, cardiovascular disease, propulsion, shoulder, handrim,
handcrank
1. Introduction
In the United States, an estimated 5.5 million peo-
ple rely on wheelchairs for their primary means of
locomotion(1) due to a variety of conditions includ-
ing congenital defects, spinal cord injuries, move-5
ment disorders, and stroke(1,2) . Though many may
associate the elderly as the most prevalent demo-
graphic among wheelchair users (WCUs), almost half
of WCUs are under the age of 65(1,3) . As mobil-
ity impairments are often permanent, WCUs depend10
on their wheelchairs and upper extremities for ambu-
lation. Unfortunately, the transition to wheelchair
use and increase in loading cycles as a result of
wheelchair-related activities may place the upper ex-
tremity at increased risk for injury(4,5,6,7,8) .15
For example, 40-75% of WCUs report upper
extremity injury and pain, often at the shoul-
∗Corresponding author: mkersh@illinois.edu (Mariana E.
Kersh)
der(9,10,11,12), which is attributed to overuse from
wheelchair propulsion and transfers(13) . This in-
creased loading of the upper arm has been suggested 20
to result in degenerative morphological changes in the
shoulder soft tissue(14), rotator cuff impingement(15),
and rotator cuff tendinopathy(6,7,16) . Even more con-
cerning is that up to 98% of WCUs demonstrate
radiographic evidence of shoulder injury including 25
asymptomatic users(16,6). In the general population,
nearly 60% of all rotator cuff injuries are asymp-
tomatic(17). The biomechanical challenges facing
WCUs are compounded by the need for increased ex-
ercise to help mitigate cardiovascular disease (CVD). 30
It is well established that physical inactivity in
WCUs increases the rates of obesity, diabetes, hyper-
tension, and dyslipidemia(18,19), which are all major
risk factors for CVD. In fact, people with spinal cord
injuries have an over 50% higher risk of CVD than 35
the general population(20). Despite clear evidence
supporting the benefits of exercise, engagement in
Preprint submitted to Engineering Archive March 25, 2022

exercise is low in WCUs due to socioecological, insti-
tutional, and interpersonal challenges(21,22). Adap-
tive sports are both physically and psychologically40
beneficial for people with spinal cord injuries (23) ,
have grown in popularity among WCU’s in recent
years(24) , and thus present an opportunity to improve
cardiovascular function. From a biomechanical per-
spective, as upper limb health is of paramount im-45
portance to WCU populations, it is essential that an
exercise intervention does not contribute to or worsen
upper limb pain.
Our understanding of shoulder function can be in-
formed by computational methods that combine in50
vivo biomechanical data (kinematics, kinetics) with
musculoskeletal models of the body. One model-
ing technique involves the use of rigid-body systems
based on subject-specific anthropometrics to simulate
different tasks using collected kinematic and kinetic55
data. The outputs of rigid-body dynamics models in-
clude joint accelerations, joint torques, muscle forces,
and joint contact forces(25,26,27,28). When combined
with computational models at the tissue level, it is
possible to obtain estimates of strains. Such models60
have been used to evaluate rotator cuff tears (29,30) ,
and have the potential to determine whether a given
exercise type may or may not place the shoulder at
risk of injury.
Critical to computational analyses of shoulder65
biomechanics during wheelchair usage is the under-
lying kinematic and kinetic data. As such, the pur-
pose of this review is to summarize existing literature
that reports continuous kinetic and kinematic data
during common locomotion modes used by WCUs.70
Within this review, we first give an introduction to
wheelchair biomechanics and commonly used terms
(Section 2) and describe our methodology for selec-
tion of papers and data reduction (Section 3). Next,
we summarized kinematic and kinetic data for two75
types of propulsion: handrim (Section 4) and crank
(Section 5). Within each propulsion type we com-
pared everyday usage to athletic forms.
2. Activities
There are two common types of propulsion modes80
used by WCU’s: handrim propulsion and crank
propulsion. Everyday wheelchair use is an example of
handrim propulsion, where the user pushes a rim on
the wheel to move forward. In contrast, crank propul-
sion (such as during handcycling), uses a gear system85
to convert rotational handcrank motion into forward
wheel motion. Within handrim and crank propulsion
are modes for athletic propulsion: wheelchair racing
and recumbent handcycling, respectively.
Figure 1: Propulsion modes: A) everyday wheelchair
propulsion, B) attach-unit handcycling, C) racing
wheelchair propulsion, D) recumbent handcycling. Angle
conventions (right-hand side) and propulsion phases for
E) handrim propulsion and F) crank propulsion. G) Force
conventions used.
2.1. Handrim Propulsion 90
Everyday propulsion features an upright
wheelchair and is the typical mode used by most
WCU’s for locomotion (Fig 1A). These wheelchairs
typically feature two large wheels, one on either side
of the user, that are perpendicular to the ground 95
along with smaller front wheels for balance. A
propulsion cycle involves grabbing the handrim and
pushing followed by a hand recovery path where no
handrim contact occurs(31). While some exercise
can be done in an upright wheelchair, WCUs switch 100
2

from an everyday wheelchair to a racing wheelchairs
to achieve faster speeds.
Racing wheelchairs also have two large wheels, but
they are placed at a camber angle for added stabil-
ity and improved speed and performance compared105
to the everyday wheelchair (32). Rather than two
smaller wheels, racing wheelchairs have one medium-
sized wheel in the front which improves the chair’s
aerodynamics and enables higher speeds (Figure 1C).
Racing wheelchairs are propelled in a crouched kneel-110
ing position, with the user leaning forward compared
to the upright sitting position of everyday wheelchair
propulsion. Instead of grabbing the handrim, propul-
sion occurs with an individual ”punching” the han-
drim, usually with a glove or hard hand-held imple-115
ment(32) ).
In both handrim propulsion modes, the kinematics
are split into two phases: the push phase, where the
hand is in contact with the wheel rim, and the recov-
ery phase, where the hand is in the air. The length120
of the push phase corresponds to the magnitude of
the contact angle defined as the end angle minus the
start angle (Fig 1E).
2.2. Crank Propulsion
Crank propulsion is an alternative mode of locomo-125
tion to wheelchair propulsion and is typically done
on a handcycle. The gears on a handcycle offer a
greater mechanical advantage than handrim propul-
sion and are suggested to create a more efficient mode
of transportation(33). Attach-unit handcycles use a130
third wheel and crank system attached to an every-
day handrim wheelchair (Fig 1B). Sport handcycles
(Fig 1D) are usually recumbent, featuring a reclined
backrest that places the body closer to the ground
in a more aerodynamic position. Handcycle propul-135
sion differs from handrim propulsion in that the hand
is in contact with the crank at all times. One cy-
cle of the handcycle crank is split into two phases:
the pull phase and push phase (Fig 1F). During the
pull phase, the user is pulling the crank toward them-140
selves; during the push phase, the user is pushing the
handle away from the body.
For the purposes of this review, we report force
data using polar conventions. Tangential force, the
force parallel to the wheel or crank path, is positive145
when in the direction of rotation. Positive radial force
was defined as the force pointing towards the center of
the wheel or center of handcycle rotation, and lateral
force, which is the force out of the saggital plane, was
defined as positive if it was pointing in the lateral 150
direction (away from the user).
3. Methods
Retrieval of papers was performed using Google
Scholar and Pubmed with combinations of the fol-
lowing keywords: ”Wheelchair”, ”Kinetics”, ”Kine- 155
matics”, ”Forces”, ”Handcycle”, and ”Inverse Dy-
namics”. Forward and backward citation searches
were used to find additional studies. Studies report-
ing kinematic and kinetic data were included if the
authors reported continuous data. Studies reporting 160
discrete (single time point) kinetic or kinematics data
were excluded. Studies reporting force data in the
global x,y,z coordinate system but not the tangential,
radial, and lateral components were also excluded. In
the 1990s, most racers shifted from an upright sitting 165
position to a forward, crouched position(31). Thus,
racing wheelchair propulsion studies published before
1990 were excluded from this review.
Continuous kinematic or kinetic data from pub-
lished figures were digitized using WebPlotDigitizer. 170
At least 15 data points were digitized per plot. Data
were processed in MATLAB and Microsoft Excel. For
handrim propulsion, only data from the push-phase
was analyzed and was normalized by length of push
phase. Joint angle descriptions are adapted from 175
International Society of Biomechanics (ISB) conven-
tions(34) for the shoulder (Fig S1).
When multiple propulsion cycles were reported per
test condition (e.g. same speed, PO, or trial), each
cycle was digitized and the cycles were averaged to 180
obtain a representative curve from the test condition.
In some cases, data from multiple studies were com-
bined using a weighted average based on the number
of study participants. All figures were rendered in R.
4. Handrim Propulsion 185
4.1. Everyday Wheelchairs: Kinetics
With the development of the SmartWheel in the
1990’s, it became feasible to record three-dimensional
handrim forces during wheelchair propulsion and pro-
vide insight into loading of the upper extremity. 190
To our knowledge, three studies have reported con-
tinuous tangential, radial, or lateral handrim force
components during everyday wheelchair use (35,36,37).
The maximum tangential forces range from 29 to 108
N (Fig 2A) and are on average greater than the range 195
3

Figure 2: Applied handrim forces (N) during everyday wheelchair propulsion (top row) and racing wheelchair propulsion
(bottom row) by A,D) tangential component, B,E) radial component, and C,F) lateral component. PP = paraplegic
subject, TP = tetraplegic subject, AB = able-bodied sub ject, WCU = wheelchair user.
of maximum radial forces (36-40 N, Fig 2B) or lateral
forces (19-33 N, Fig 2C). Tangential forces have a sin-
gle peak in the last half of the push phase (54-79%)
whereas radial and lateral forces have multiple peaks.
Radial forces tend to reach a maximum early in the200
push phase (14-22%). In contrast, lateral forces tend
to have two peaks: an initial lower peak in the first
half of the push phase (12-27%) followed by a higher
peak in the second half of the push phase (63-78%).
Differences in the magnitude of the tangential205
forces may be due to different wheelchair setups and
speeds used during testing. For example, the largest
reported tangential forces occurred with subjects pro-
pelling at 60-80% of peak power outputs(36) while the
other force profiles likely occurred at lower speeds210
(range: 1.5-2.0 m/s) (37,35) . Experienced wheelchair
users tended to apply lower peak tangential forces
(p=0.0001) and took longer to reach the peak tan-
gential forces (p=0.0015) than able-bodied partic-
ipants(37) , though one exception is reported from215
Dallmeijer and colleagues(36).
Notably, the lateral force component reported from
a subject with tetraplegia (36) was negative and with
a single peak while others report positive double-
peaked lateral forces. Whether or not the location 220
and severity of spinal cord injury affects the loads ap-
plied during propulsion remains to be clearly demon-
strated, but these data suggest that care should be
taken when interpreting wheelchair kinetics from par-
ticipants of varying injury levels. 225
4.2. Wheelchair Racing: Kinetics
An examination of the studies that report contin-
uous applied forces during racing wheelchair propul-
sion yields results with high variability in both the
shape of the force curves and the location of peak 230
forces (Fig 2D-F). In contrast to everyday wheelchair
usage, the maximum forces were applied radially
during racing wheelchair propulsion. As expected
due to increased speeds, the average maximum ap-
plied forces during racing propulsion were larger for 235
all magnitudes of components compared to everyday
propulsion (118 vs. 64 N for tangential forces, 251 vs.
38 N for radial forces, and 94 vs 26 N for lateral forces
in racing vs. everyday propulsion, respectively). Sim-
ilar to everyday wheelchair propulsion, the maximum 240
tangential forces (range: 105-131 N, Fig 2D) during
racing propulsion peaks once in the middle of the
4

Figure 3: Joint angles (°) during the push phase of everyday wheelchair propulsion, averages in blue.(38,39)
push phase (40-69%) compared to the latter part of
the push phase during wheelchair propulsion.
Also similar to everyday, radial forces during racing245
peaked multiple times with maximum forces ranging
from 150-428 N, but all force profiles peaked in the
second half of the push phase (64-82%) compared to
earlier peaks at the beginning of the push phase for
everyday propulsion (Fig 2E). Chenier et al. reported250
maximum radial forces of 427 N, which they attribute
to the design of the force measurement system and
the fact that their participant was propelling at max-
imum speed(40). In contrast to everyday propulsion,
two of the three racing studies reported negative lat-255
eral forces but with inconsistent force profiles. The
positive lateral forces reported by Limroongreungrat
et al. were attributed to differences in wheelchair de-
sign and propulsion speed(41).
4.3. Everyday Wheelchairs: Kinematics260
Kinematic profiles of the shoulder within the two
studies reporting shoulder angles using the ISB con-
ventions were consistent. Wheelchair propulsion be-
gins with the shoulder abducted to 53o, flexed to 48o,
and externally rotated by -83o(Fig 3A-C, respec-265
tively). During the push phase, the shoulder adducts
to 30o, extends to -12o, and internally rotates to -19o.
Boninger et al.(42) and Koontz et al.(43) also re-
ported continuous shoulder angles, but these studies
were done prior to ISB shoulder angle conventions270
and used projections of the humerus in the anatom-
ical planes. As a result, these shoulder angles were
not included in our analysis but are worth noting as
sources of continuous kinematic data.
4.4. Wheelchair Racing: Kinematics 275
Few studies report continuous joint angle data dur-
ing racing propulsion using the current (post-1990’s)
racing wheelchair design. To our knowledge, two
studies have reported continuous joint angles, but
only reported elbow flexion(44,45). Moss et al. (46) ,280
who reported start and end angles of a sprint start
during wheelchair racing, was also excluded from
analysis due to the altered technique associated with
static propulsion mechanics.
There are however reports of the start and end an- 285
gles during propulsion (Fig 4). These contact angles
are important to quantify because the length of the
push phase affects the length of time that the shoul-
der experiences applied forces(47). In racing propul-
sion, the mean start angle is 27.8°and the mean end 290
release angle is 197°, on average (Fig 4). The result-
ing contact angle of 169°indicates that the hand is
in contact with the handrim and transferring forces
for almost half of the propulsion cycle. In everyday
propulsion, the average push angle is 83°.295
Start angles and end angles vary depending on the
handrim propulsion mode, with everyday propulsion
having a smaller and more precise push phase com-
pared to racing propulsion. The increased variability
in the push phase for racing propulsion could be at- 300
tributed to racing technique (recovery path location,
having a tightly closed fist or using the thumb to ap-
ply force, etc.) based on personal preference and race
length(48,49).
5

Figure 4: Handrim contact and release angles for
racing(50,49,48,51,52,53,46,44,47) and everyday(36,43,54,55,56,57)
propulsion. Mean angles for each activity are shown in
black.
4.5. Implications of handrim wheelchair vs racing305
propulsion
In general, the maximum applied force compo-
nents during racing wheelchair propulsion are 1.2-
10.7 times larger than those recorded during every-
day wheelchair propulsion. This increase is likely due310
the high speeds and intensities associated with rac-
ing. Whether these increased force values may be
placing the shoulder at increased risk for injury re-
mains to be clearly demonstrated, though Mercer et
al. have shown that increased lateral forces are as-315
sociated with pain and injury(58). The wide range
of force values and profiles across studies points to
the need to more fully examine the applied forces of
racing propulsion including larger sample sizes with a
more consistent method of measuring applied forces.320
The lack of shoulder kinematic data during racing
prevents a direct comparison to everyday wheelchair
usage. However, the start and end angles can also
lend insight into changes in propulsion styles. While
the longer (2 times the contact angle) push phase in325
racing compared to everyday propulsion may indicate
more time for force transfer, and therefore less impact
on the shoulder, racing propulsion also results in in-
creased forces. It is still unclear how the combination
of increased forces and larger contact angle during330
racing propulsion affects shoulder joint loading, and
therefore tissue strain and injury risk.
The force transfer during wheelchair propulsion is
also dependent on propulsion kinematics: increased
shoulder elevation, and shoulder rotation are corre- 335
lated with increased shoulder joint reaction forces
during racing propulsion, which is indicative of higher
shoulder injury risk(47). Thus, it is important to more
carefully quantify how the upper arms are moving
during wheelchair racing propulsion to optimize force 340
transfer from an injury prevention standpoint.
5. Crank Propulsion
With the introduction of a crank for propulsion,
applied forces and movement shift from the period
associated with the push angle of propulsion to con- 345
tinuous biomechanical data throughout a complete
propulsion cycle. One should note that the origin
for the angular phases during crank propulsion is dif-
ferent (anterior) than the origin used for handrim
propulsion (proximal). 350
5.1. Attach-unit Handcycling: Kinetics
Two studies have reported continuous forces dur-
ing attach-unit handcycling, with only one study(59)
reporting all three force components. The tangen-
tial force profiles have similar shapes and primarily 355
differ in the magnitude of the forces (Fig 5). The
largest peak tangential force recorded was 45 New-
tons and occurred between 64°and 91°of the cycle
(Fig 5). The transition from push to pull phase is
indicated by the local minimums of tangential forces 360
which occurred between 276°and 310°. As rolling re-
sistance increased, shown via increases in Watts, tan-
gential forces also increased to maintain speed(60) .
Able-bodied subjects cycling at 1.94 m/s had lower
tangential forces(60) than those reported in a subject 365
with paraplegia cycling at 35W(59).
5.2. Recumbent Handcycling: Kinetics
To our knowledge, only one study has reported di-
rect measurements of continuous applied forces dur-
ing recumbent handcycling (Fig 6A)(66) . Ahlers and
Jakobsen report a maximum tangential force of 127
N at 107°in the propulsion cycle, with a pronounced
push-pull and pull-push transition at 2°and 199°, re-
spectively. Several studies have investigated the ef-
fect of changing either the handcycle configuration
(i.e. crank length, backrest angle, crank position,
etc.) or the power output during propulsion and re-
ported continuous torque during recumbent handcy-
6

Figure 5: Tangential force (N) during attach-unit hand-
cycling(59,60). The three shades of blue represent three
resistance levels, but at varying gear levels.
cling(64,65,62,63). This torque data can be used to in-
directly calculate Ftan using Equation 1:
Ftan =τ/R (1)
where τis the torque at the crank and R is the crank
length. Changes in crank length do not significantly
change the applied torque profiles(62) (Fig 6B, blue370
lines). However, torques are sensitive to the crank po-
sition(63): when the crank is moved closer to the par-
ticipant, the torque in the push phase increases and
torque in the pull phase decreases (Fig 6B). The tan-
gential forces during recumbent handcycling kinetics375
are also sensitive to changes in power output with
maximum force increasing with increasing power.
The location of maximum force across studies was
variable (range = 34°- 273°), and could be a result
of varying levels of handcycling experience, handcy-380
cle design, participant demographics, and different
methodologies in collecting force data. The loca-
tion of the push-pull transition (291°-2°) and pull-
push transition (132°-199°) (Fig 1F), defined as the
location of minimum applied force, were more consis-385
tent across studies though studies investigating the
effect of power (Fig 6C) did not report a clear tran-
sition point. Compared to attach unit handcycling,
the push-pull transition occurs later in the propulsion
cycle. It is worth noting that the study directly mea-390
suring Ftan (Fig 6A) has a force profile similar to that
recorded during attach unit handcycling, with the
maximum tangential force occurring at the bottom of
the propulsion cycle (Fig 5). It’s possible that, with
more consistent data collection methods across re- 395
cumbent and attach-unit handcycling, a clearer tan-
gential force curve would emerge.
5.3. Attach-unit Handcycling: Kinematics
We did not find any studies that reported con-
tinuous joint angles during attach-unit handcycling. 400
Faupin and Gorce reported the maximum and mini-
mum shoulder angles for one able-bodied participant
and one participant with paraplegia during 70 rpm
handcycling(70). They reported a 63, 19, and 11 de-
gree range of motion for shoulder flexion, abduction, 405
and rotation, respectively, for the able-bodied par-
ticipant. The range of motion for the subject with
paraplegia was greater than the able-bodied subject,
with 71, 23, and 17 degree ranges of motion for shoul-
der flexion, abduction, and rotation, respectively. 410
5.4. Recumbent Handcycling: Kinematics
In contrast to attach-unit handcycling, continuous
kinematic data during recumbent handcycling has
been reported. Handcycling begins with the shoul-
der in abduction after which it continues to abduct 415
by 31°, followed by adduction to 10°(Fig 7A). Av-
erage maximum abduction of 31°occurs at 199°in the
propulsion cycle. There is a slight increase in shoulder
abduction with an increase in power output: as the
intensity of the exercise increased, the shoulder be- 420
comes more elevated and abducts up to 39°. Shoulder
flexion exhibits two peaks during recumbent handcy-
cling: initially the shoulder is flexed to X°and then
extends to -19°followed by a return to flexion which
peaks at 300°towards the end of the propulsion cycle 425
(Fig 7B).
The variation between studies in shoulder abduc-
tion and flexion was less than the variation in shoul-
der internal / external rotation (Fig 7C). Studies that
used acromion marker clusters(71) to track scapular 430
movement reported a wider variation in shoulder ro-
tation, from -10°to 45°(68,62). The studies that did
not use an acromion cluster to track scapula kine-
matics reported a smaller range of shoulder rotation,
from about 10°to 25°(64,67).435
5.5. Implications for attach unit vs recumbent hand-
cycling
During sprinting, the tangential applied forces
reach a maximum of 447 N, which is more than ten
7

Figure 6: Tangential applied force (Ftan) (N) during recumbent handcycling. Studies examining A) Ftan measured directly
using a strain gauge-instrumented handle, self-selected speed(61), B) changes in handcycle configuration (changes in crank
length shown in blue(62) , and changes in crank fore-aft position, as a percentage of arm length, shown in orange and
green(63)) and C) changes in power (20-120W)(64) and sprinting(65).
times larger than the maximum tangential forces re-440
ported during attach-unit handcycling. When ath-
letes train at increased speeds and power outputs, it
is reasonable to suggest that these increased external
forces will result in increased loads at the shoulder.
The question remains as to which power or speed lev-445
els, and for how long, are acceptable to avoid overuse
injuries common in wheelchair athletes. Additionally,
only continuous Ftan data was available in the liter-
ature for recumbent handcycling. While Ftan is the
majority of the total applied force during attach-unit450
handcycling, it’s still unclear how the other compo-
nents (radial and lateral) are impacted by changes
in recumbent handcycle configurations and power.
Force changes for all three directions are important
for modeling the upper extremities and understand-455
ing musculoskeletal loads during any activity includ-
ing recumbent handcycling(72) .
Kinematics differences are also important: the ele-
vated positions associated with increased power have
been suggested as a source of increased risk for in- 460
jury(73). Because of the lowered body position in
recumbent handcycling, participants are propelling
with their arms at more elevated positions compared
to attach-unit handcycling. Other overhead sport
activities (i.e., swimming, volleyball, gymnastics), 465
which also involve elevated shoulder angles, are con-
sidered to put the shoulder at more injury-prone po-
sitions(73). This is consistent with findings by Ar-
net et al. who confirmed that a more inclined back-
rest position (more recumbent) leads to higher shoul- 470
der loads in attach-unit handcycling(74) . While the
Figure 7: Recumbent handcycling joint angles (°), averages in blue. (67,62,64,65,68,69)
8

impact of shoulder rotation on shoulder soft tissue
strain hasn’t been quantified for recumbent handcy-
cling, higher internal rotation was identified as a risk
factor during weight bearing activities for shoulder475
pain in WCU’s(75) . Nonetheless, more data is needed
to identify potential stages during crank propulsion
when the load on the shoulder is highest to optimize
training protocols to reduce injury risk.
While this is a promising start, a more thorough480
investigation of the three-dimensional shoulder loads
during recumbent handcycling using models specific
to WCUs could provide insight into injury risk and
prevention techniques, especially at the higher speeds
experiences during recumbent handcycling races and485
exercise.
6. Conclusion
Upper extremity injuries and CVD in WCUs rep-
resent a multi-factorial problem with an evidenced
burden on individuals and society. A greater under-490
standing of upper-limb kinetics and kinematics dur-
ing propulsion can lend insight into joint loading,
and therefore joint forces and injury risk, faced by
WCUs during both handrim and crank propulsion.
In general, the applied hand forces are larger for both495
athletic propulsion modes (racing wheelchair propul-
sion and recumbent handcycling). These increased
forces could indicate increased shoulder loading dur-
ing these activities. While this data is a promising
start to characterizing the biomechanics of common500
propulsion modes in WCUs, there is still work to be
done. With continuous kinetic and kinematic data,
upper limb rigid-dynamics models can be created that
calculate joint torques and accelerations, joint con-
tact forces, and muscle forces, which can lend in-505
sight into shoulder injury risk. There is still a need
to more completely characterize shoulder joint angles
during racing wheelchair propulsion and continuous
three-dimensional forces during recumbent handcy-
cling. Once these propulsion styles are more com-510
pletely characterized, we can develop predictive mod-
els for coaching, training, and physical therapy to re-
duce shoulder injury risk and increase the quality of
life for WCUs.
9

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12

Supplemental Data
13

Figure S1: Joint angle conventions based on ISB standards for A) shoulder abduction / adduction, B) shoulder flexion
/ extension in the sagittal plane, and C) shoulder internal / external rotation. Angles are defined as the angle of the
humerus relative to the thorax.
14